Field of the Invention
The present invention concerns a method to correct a signal phase in the acquisition of MR signals of an examination subject in a slice multiplexing method, and an MR system for implementing such a method.
Description of the Prior Art
The desire for ever faster MR acquisitions in the clinical environment is currently leading to a renaissance of methods in which multiple images are acquired simultaneously. In general, these methods can be characterized in that transverse magnetization of at least two slices is specifically used simultaneously for the imaging process (“multislice imaging”, “slice multiplexing”), at least during a portion of the measurement. In contrast to this, in the established “multislice imaging” the signal is acquired from at least two slices in alternation, i.e. completely independently of one another with correspondingly longer measurement time. For example, the following are among such slice multiplexing methods:
Hadamard coding (for example Souza et al., J. CAT 12:1026 (1988)): two (or more) slices are excited simultaneously, a defined signal phase is impressed on each slice via corresponding design of the RF excitation pulses. The signal of the magnetization is received simultaneously from both slices. A similar second excitation of both slices is implemented, but with different relative signal phase in the slices. The remaining imaging process (phase coding steps) takes place in a conventional manner; the method can be combined with any acquisition techniques ((multi-)gradient echo, (multi-) spin echo etc.). The signal information of both slices can be separated from the two exposures by means of suitable computing operations.
Simultaneous echo refocusing (SER, SIR, for example Feinberg et al., MRM 48:1 (2002)): two (or more) slices are excited simultaneously. The signal of the magnetization is received simultaneously from both slices. During the data acquisition, a gradient is activated along the slice normals, which leads to a separation of the signals of both slices in frequency space. The remaining imaging process (phase coding steps) takes place in a conventional manner; the method can be combined with any acquisition techniques ((multi-)gradient echo, (multi-) spin echo etc.). Images of both slices can be separated from the simultaneously acquired data by means of suitable computing operations.
Broadband data acquisition (for example Wu et al., Proc. ISMRM 2009:2768): two (or more) slices are excited simultaneously. The signal of the magnetization is received simultaneously from both slices. During the data acquisition, a gradient is activated along the slice normals, which leads to a separation of the signals of both slices in frequency space. The remaining imaging process (phase coding steps) takes place in a conventional manner; the method can be combined with any acquisition techniques ((multi-)gradient echo, (multi-) spin echo etc.). The signals of both slices can be separated from the simultaneously acquired data by means of suitable filtering.
Parallel imaging in the slice direction (for example Larkman et al., JMRI 13:313 (2001)): two (or more) slices are excited simultaneously. The signal of the magnetization is received simultaneously from both slices with at least two (or more) coil elements. The remaining imaging process (phase coding steps) takes place in a conventional manner; the method can be combined with any acquisition techniques ((multi-)gradient echo, (multi-) spin echo etc.). An additional calibration measurement is implemented to determine the spatial acquisition characteristic of the coil elements. The signals of both slices can be separated from the simultaneously acquired data by means of suitable computer operations (GRAPPA algorithm, for example).
Furthermore, in single slice imaging it may be necessary to correct image artifacts given which the correction parameters depend strongly on the spatial position or on the signal of the individual slices. An example of this is the correction of phase errors that arise due to accompanying Maxwell fields. These phase errors arise in that there is no complete linearity of the magnetic field gradient upon switching of a linear magnetic field gradient; rather higher-order terms always arise. These fields—known as Maxwell fields—lead to phase errors in the detected MR signals. One possibility for correction is described in Meier et al., MRM 60:128 (2008). Likewise, in single slice imaging it is sometimes necessary to correct local inhomogeneities of the basic magnetic field that would lead to signal cancellations or image distortions. The correction of such inhomogeneities is described in Deng et al., MRM 61:255 (2009) and in Lu et al., MRM 62:66 (2009), for example.
In many cases of slice-specific correction in single-slice exposures it is sufficient to merely impress an additional linear signal phase along the slice coding. Using various examples it is subsequently explained why the impression of a linear signal phase is sufficient in many cases:
a) One possibility of application of a linear correction is the correction of phase errors in diffusion imaging that are due to Maxwell fields.
MRM 60:128 (2008) describes how the accompanying fields of the Maxwell fields of the diffusion coding gradients lead to an additional signal dephasing along the three spatial coordinate axes. Dephasings along the frequency and phase coding axis merely lead to a displacement of the signal in k-space—the echo is no longer acquired at k=0, but rather at a (slightly) shifted position. An echo shift in k-space corresponds in positional space (after the Fourier transformation) to a linear phase response in the image; insofar as only magnitude images are of interest, this effect only plays a subordinate role. Moreover, by the acquisition of a sufficiently large k-space region (omitting partial Fourier techniques, for example) it can be ensured that the echo signal is located in the scanned region in every case.
However, dephasing along the slice selection axis directly leads to a signal loss that cannot be compensated. The magnitude of the dephasing thereby depends on the amplitude of the accompanying Maxwell fields (and thus on the position of the slice). Given simultaneous acquisition of multiple slices, an individual dephasing is to be corrected for every slice. In the first order, the dephasing can be described by a linear phase response.
b) A linear phase correction is likewise sufficient given the correction of the phase errors of flow imaging that are due to the Maxwell fields.
As in the preceding example, this example relates to the compensation of dephasings due to accompanying Maxwell fields—here caused by the gradients used for the flow coding. The statements regarding the motivation of a linear, slice-specific correction phase along the slice coding axis analogously apply to this example. The uncorrected linear phase response in the image (due to the shift of the echo in k-space) can be taken into account in the data processing in a simple manner.
c) A linear correction is likewise possible to correct local inhomogeneities of the basic magnetic field and the signal cancellations that are caused by these (z-shim).
MRM 61:255 (2009) (and the references cited therein—in particular in Yang et al., MRM 39:402 (1998)) describe how imaging errors in echoplanar gradient echo imaging that are caused by inhomogeneities of the basic magnetic field can be reduced by repeated implementation of the measurement with different auxiliary gradients in the slice coding direction. These are (local) magnetic field gradients that lead to a dephasing of the signal along the three spatial coordinate axes. Again, it is only the gradient along the slice coding direction that has the largest effect on the image quality due to the signal loss within a voxel (intra-voxel dephasing) associated with this gradient direction. The known z-shim method varies a background gradient from measurement to measurement in order to ensure a good rephasing of each slice in at least one measurement for each every spatial region. The multiple images of a slice are merged into an image with reduced signal cancellations, either by a simple averaging (absolute mean value, “sum of squares”) or by more complicated combination methods.
d) Linear phase correction can likewise be applied in the correction of signal cancellations and image distortions that result due to local inhomogeneities of the basic magnetic field (SEMAC).
MRM 62:66 (2009) describes how signal cancellations and image distortions in 2D imaging that are due to metal implants (or, respectively, the local inhomogeneities of the basic magnetic field that are connected with these) can be reduced via use of a (limited) additional phase coding along the slice normal. Similar to the case of a z-shim, multiple measurements with different auxiliary gradients in the slice coding direction are acquired per slice, and these data are combined in a suitable manner.